The present invention relates to magnetic resonance imaging (MRI) and, more particularly, to an MRI system capable of high-speed imaging processing for obtaining an instantaneous image of a region including a dynamic object such as a heart.
As is well known, an MRI method utilizes a magnetic resonance (MR) phenomenon to two-dimensionally or three-dimensionally obtain chemical and/or physical microscopic information of a material. The MR phenomenon is a phenomenon wherein when a group of nuclear spins, which are unique to the types of atomic nuclei and have magnetic moments, are placed in a uniform static magnetic field, an energy of a radio-frequency (RF) magnetic field, which rotates at a specific frequency in a plane perpendicular to a direction of the static magnetic field, is resonantly absorbed.
In the MRI, the spatial distribution of specific atomic nuclei in an object to be examined (typically, hydrogen atomic nuclei in water or fat in a body when an object is a living body such as a human body or an animal) is imaged.
As known MRI methods, the projection reconstruction method by Lauterbur, the Fourier method by Kumar, Welti, and Ernst, the spin warp method by Hutchison et al. as a modification of the Fourier method, the echo planar method by Mansfield, and so on, have been proposed.
In an MRI system, in order to acquire data necessary for reconstructing an image, various magnetic fields are combined as needed, and are applied to an object to be examined in accordance with a predetermined sequence. In this case, data acquisition must be performed such that predetermined magnetic field application sequences are repeated while changing the intensity and/or application time of a gradient field in a specific direction of the magnetic fields applied to the object. Therefore, a data acquisition operation, i.e., a scan operation, for long periods is necessary. For this reason, it is not easy to obtain an instantaneous image of a region including a dynamic object, such as a heart, whose position and/or shape changes over time. When such an image is obtained, an image blurring and/or artifact inevitably occurs as long as a special method such as a cardiac cycle synchronization method is not adopted. In the cardiac cycle synchronization method, data acquisition is performed for short periods in synchronism with the movement of a dynamic portion at timings at which the dynamic portion is in the same state. In the cardiac cycle synchronization method, data acquisition is intermittently performed for short periods with relatively long rest periods. For this reason, an imaging time required for obtaining all the necessary data is prolonged. The method, such as the cardiac cycle synchronization method, for acquiring data in synchronism with the movement of the object has limited objects to be applied or to be observed, and cannot be applied to diagnosis of the heart function by time-serially observing the movement of the heart.
As methods for solving the above problem, high-speed imaging methods such as the echo planar method or the fast Fourier (FF) method have been proposed. In these methods, it was demonstrated that an MR image of the internal organs of a human body could be obtained in a short time, e.g., in about 50 msec.
The high-speed imaging methods include the FID (free induction decay) method for observing an FID signal and the spin echo method for observing a spin echo signal. The FID method and the spin echo method have different ways to excite the MR phenomenon. In this case, the spin echo method will be described below.
FIGS. 1A to 1E show pulse sequences of the spin echo method in the conventional echo planar method and the FF method.
In the echo planar method, gradient field Gz in a z direction for selecting a slice to be imaged is applied (FIG. 1B), and a 90.degree. pulse (an excitation pulse, whose flip angle of the magnetization vector caused by excitation is 90.degree., is called a 90.degree. pulse) is applied while gradient field Gz is applied (FIG. 1A). After application of the 90.degree. pulse, gradient field Gz is inverted (FIG. 1B). After application of inverted gradient field Gz, gradient field Gy in a y direction (perpendicular to the z direction) is applied for a predetermined period of time (FIG. 1D). Thereafter, gradient field Gz is applied (FIG. 1B), and a 180.degree. pulse (an excitation pulse with which a magnetization vector of a nuclear spin is inclined through 180.degree. as a result of excitation) is applied while gradient field Gz is applied (FIG. 1A). After a predetermined period of time has passed from the application of the 180.degree. pulse, gradient field Gx in an x direction (perpendicular to magnetic fields in the y and z directions) is applied (FIG. 1C), and gradient field Gy is applied to be superimposed on gradient field Gx (FIG. 1D). Gradient field Gx is repetitively inverted at predetermined timings while gradient field Gy is applied (FIG. 1C).
In the FF method, the excitation pulses and gradient fields Gz and Gx are applied in the same manner as in the echo planar method, as shown in FIGS. 1A to 1C. However, gradient field Gy is applied differently from the echo planar method, as shown in FIG. 1E. In this method, pulsed gradient field Gy is applied upon second inversion and thereafter of gradient field Gx.
The conventional high-speed imaging methods include the rapid projection method and the spiral scanning method in addition to the above-mentioned echo planar method and the FF method. FIGS. 2A to 2D show the pulse sequences of the rapid projection method in the spin echo method, and FIGS. 2A, 2B, 2E, and 2F show the pulse sequences of the spiral scanning method.
In the conventional high-speed imaging methods, the gradient field must be repetitively inverted at high speed, so as to generate spin echoes a predetermined number of times determined by an image matrix. For example, if a matrix size of an image is (2N.times.2N), (N+1) echo signals must be acquired. Therefore, if N=32, 33 echo signals must be acquired. In this case, the gradient field must be repetitively inverted at a very high speed in consideration of nonuniformity of the gradient field or offset from resonant point. In addition, since the gradient field must have a very high intensity (i.e., the degree of gradient), this may cause an adverse influence on an object to be examined if the object is a living body.
When nonuniformity .DELTA.H(x,y) of the static field is present (assuming the case wherein an object is sliced along a plane perpendicular to the z direction), a transversal magnetization vector of point (x,y) is subjected to phase modulation at an angular velocity of .gamma..DELTA.H(x,y). Therefore, a phase error of the transversal magnetization is accumulated proportional to an elapsed time from application of the RF pulse for exciting the MR phenomenon. For example, phase error .phi.n of the nth echo signal is given by .phi.n=.gamma..DELTA.H(x,y)nTI (TI is a time interval between inversions of the gradient field). Assuming that the nonuniformity of the static magnetic field is .vertline..DELTA.H(x,y).vertline. max/H0 (H0 is the central static magnetic field intensity) and the intensity of the static magnetic field is 0.5T, .vertline..gamma..DELTA.H(x,y).vertline. max is about 40 .pi.. In this case, if the phase error is to be suppressed to .pi./10, NTI.apprxeq.2.5 msec, and if N=32, TI=80 .mu.sec. In this case, the inversion of the gradient field must be completed within a time of about 10 .mu.sec. However, it is impossible to invert the gradient field in such a short period of time for the purpose of safety of an object (patient) and in terms of technique. This also applies to a case wherein the inversion of the gradient field is performed in accordance with a moderate waveform such as a sine wave, like in the spiral scanning method.
The gradient intensity when the gradient field is inverted according to a rectangular wave will be evaluated. If a one-dimensional length (diameter) of an object portion to be imaged is 20 cm, and an image matrix is (64.times.64), spatial resolution .DELTA.l of 3 mm can be obtained. Since a frequency resolution in observation time TI=80 .mu.sec is .DELTA.f=1/TI=12.5 kHz, the gradient field is G=.DELTA.f/.DELTA.l=42 kHz/cm. This value is about 5 times the maximum gradient intensity of 8 kHz/cm used in the conventional system for a human body, and may cause an adverse influence on the human body. In addition, this value is difficult to attain in terms of the present technique. Assuming that inversion time .DELTA.TI of the gradient field (a time required for completing inversion) is 10 .mu.sec, a changing ratio of the gradient field is G'=2G/.DELTA.T.apprxeq.10.sup.10 Hz/cm.multidot.sec, and this may cause an adverse influence on the human body. If static magnetic field intensity H0 is decreased, the abovementioned requirements for the inversion time and intensity of the gradient field are moderated since .DELTA.H(x,y).infin. H0. However, if H0= 0.1T, this value is still difficult to attain.
A signal to be observed is given by the spatial integral associated with x and y of magnetization present at point (x,y). For this reason, if the phase error is accumulated due to the influence of the nonuniformity of the magnetic field, as described above, a complicated distortion from correct Fourier data occurs. Since this distortion is not simple, the influence of the nonuniformity of the magnetic field cannot be removed by computational processing of a known algorithm for the Fourier method and the projection method. The influence of the nonuniformity of the magnetic field in the high-speed imaging not only causes a mere spatial distortion of an image, but also causes blurring and a noticeable artifact of an image.
In this manner, since it is difficult to compensate for the influence of the nonuniformity of the magnetic field by software processing, it is required to reduce the nonuniformity itself. However, it is technically difficult to reduce .DELTA.H(x,y) to a negligible level. As the magnetic field intensity is increased, .DELTA.H(x,y) increases proportionally thereto. Therefore, the abovementioned difficulty is still enhanced.
For these reasons, the conventional high-speed imaging methods can only realize to reconstruct a relatively small object in a low magnetic field of about 0.1T. An image having an artifact can only be obtained by imaging under these conditions. If the low magnetic field is used, an S/N (signal-to-noise) ratio is further degraded in addition to poor S/N ratio as the nature of high-speed imaging, and the image quality is considerably degraded. For these reasons, although high-speed imaging has excellent features, it cannot be applied to actual clinical examination.
In order to eliminate an influence of the nonuniformity of the magnetic field and offset of the magnetic field intensity from the resonance point and to allow an algorithm for compensating for an image distortion caused by the nonuniformity of the magnetic field to be applied, it can be considered that an echo signal operation by applying a 180.degree. pulse is used instead of that by repetitively applying the gradient field in the conventional high-speed imaging. More specifically, an operation for applying a 180.degree. pulse a plurality of number of times to produce a multiple echo signal and an operation for applying predetermined magnetic fields are combined based on the CPMG (Carr Purcell Meiboom Gill) method or the modified CP (Carr Purcell) method known in the NMR spectroscopy, and entire image reconstruction data is acquired by single excitation using the 90.degree. pulse. In this case, 180.degree. pulse application is performed so as not to accumulate the phase error of the echo signal.
However, in the CPMG method and the modified CP method, a very large number of 180.degree. pulses must be applied. For this reason, if an object is a human body, the above methods may adversely influence the human body. The 180.degree. pulses used can include selective excitation pulses for exciting only magnetization in a specific slice plane and nonselective excitation pulses for exciting the entire predetermined imaging region. When the selective excitation pulses are used, a necessary power can be reduced as compared to the case of the non-selective excitation pulses, and multi-slice imaging (for time sharing imaging a plurality of different slices during an imaging time for a single slice utilizing a magnetization recovery period of time) can be allowed. However, as the number of echoes of a multiple echo signal string is increased caused by problems associated with the apparatus arrangement, a slice characteristic is degraded, and an echo signal level is decreased upon degradation. As a result, this causes a hazard to the human body, and an image quality of a resultant MR image is degraded. When the nonselective excitation pulses are used, this may cause a serious hazard to the human body. In this case, multi-slice imaging cannot be performed. In addition, flip angles of magnetization caused by excitation pulses may be shifted, and a pseudo FID signal as an FID signal which would not be produced after the 180.degree. pulse may be produced due to the nonuniformity of the 180.degree. pulses. Therefore, an artifact may be formed on an MR image due to mixture of the pseudo FID signal.
In this manner, the conventional high-speed imaging methods may cause image blurring due to the nonuniformity of the static magnetic field, generation of artifacts, and degradation in S/N ratio, at a practical static field intensity. In the CPMG method or the modified CP method which can eliminate the influence of the nonuniformity of the static magnetic field, this may cause a serious hazard to an object to be examined.